Magnetic resonance imaging is a well established medical diagnostic method that achieves good resolution, especially at main magnetic fields of 0.5 Tesla and above. The main magnetic field is a static field which should be stable and homogeneous. In the majority of systems superconducting magnets are used but there are many systems that use resistive and permanent magnets especially for fields lower than 0.5 Tesla.
Since the resonance frequency of the nuclei is a function of the magnetic field strength, MRI techniques use field gradients in the magnetic field in order to provide a unique code for every point in the imaging volume. The gradient coils used to generate the field gradients usually have the form of cylinder for the Z direction gradient (axial) coil and a saddle or "stream function" form for the X and Y (transverse) direction gradient coils.
The gradient coils are energised by pulses. The changing magnetic field of the gradient coils induces currents in any proximate electrically conductive material. These induced currents or eddy currents distort the magnetic field in the imaging volume in both space and time, and consequently adversely affect the quality of the images.
One of the parameters that determines the cost and imaging ability of the main magnet system is the free bore diameter. Reducing the diameter of the free bore of the magnet increases the efficiency of the magnet in terms of cost and field strength. This is true for the magnets, whether they are super-conducting, resistive or permanent magnets.
One big benefit in reducing the diameter of the free bore of the magnet is that the volume enclosed in the 5 Gauss line (used to measure the magnetic interference volume) is decreased as well. Accordingly, problems associated with shielding and safety are much easier to overcome. The main drawback when reducing the diameter of the free bore is that there is a consequent increase in the eddy currents which adversely affect the magnetic field in the imaging volume.
In the prior art, there are two main solutions to the eddy current problems. The first is the use of a shield of conducting material which is placed between the gradient coil and the magnet. The shield protects the magnet system, but eddy currents are induced in this shield which still distort the main field in many instances, depending on the pulse sequence.
The second method is the use of an additional set of gradient coils which has a bigger diameter than the first set of gradient coils. The first set of coils produce a gradient field in the imaging volume, but also produce a magnetic field outside the coil. The second set of coils produces a field outside the coils which is equal to the field produced by the first set but in the opposite direction. The net result is that outside the gradient system the gradient field is very low. However, the field strength in the imaging volume is also reduced. In order to get the required field strength in the imaging volume the power in both sets of coils is increased.
A drawback with this approach is that in order to achieve good shielding, the radius of the outer set of coils needs to be much bigger than the radius of the first set of coils, and the free bore of the magnet cannot be reduced. Another drawback is an increase in the power required for the same gradient in the center of the imaging volume.